CROSS REFERENCE TO RELATED APPLICATIONS

[0001] This application claims priority pursuant to 35 U.S.C. § 119(e) to U.S. Provisional Patent Application Ser. No. 62/385,592, filed on September 9, 2016, which is hereby incorporated by reference in its entirety.

FIELD

[0002] The present invention relates generally to sensors, sensor systems, and methods for making and using the same. Particularly, the present invention relates to redox sensors exhibiting improved sensitivity, sensor systems using the same, and methods for making and using the same.

BACKGROUND

[0003] Redox sensors, which comprise a redox catalyst, have a wide variety of applications, including industrial applications, food processing applications, clinical medicine applications, and athletic performance monitoring.

[0004] Enzyme sensors, where the enzyme is the redox catalyst, are one known type of redox sensor, and Lactate enzyme sensors are one widely used type of enzyme sensor. Improved lactate detection by an in situ sensor is of great need in the areas of clinical medicine, food processing, and athletic performance monitoring.

[0005] Lactate is a common analyte due to its wide variety of applications. In the field of food processing, L-lactate is present in the fermentation of cheese, yoghurt, butter, pickles, sauerkraut, and other food products. Monitoring lactate concentrations can be used to assess the condition of freshness in dairy products. In fish farming, health and stress levels of fish can be monitored by testing the lactic acid concentrations, thereby limiting the mass use of antibiotics and the inadvertent consequences byhuman consumption of the residual antibiotics. In human bodies, lactate is metabolized predominantly in the kidney and liver. The normal lactate level range in the human body is 0.5-2.0 mM/L. Hyperlactatemia is defined when lactate levels are between 2 mM/L (millimol/liter) to 5 mM/L. When lactate levels exceed 5 mM/L, the conditions indicate severe lactic acidosis. In the fields of human performance monitoring, lactate levels play a key role to brain blood flow and have an impact on brain activation during exercise-induced fatigue. Transient lactic acidosis can occur due to excessive lactate production from tissue hypoxia or increased cellular metabolism caused by strenuous exercise. By monitoring the lactate thresholds in endurance athletes, a lactate sensor can inform users of their limits during exercise. In clinical medicine, hyperlactataemia is an indicator of systemic tissue dysoxia or abnormal microcirculatory perfusion. It may also indicate the severity liver injury and the accompanying multiple organ failure. Lactic acidosis is indicative of tissue ischemia, liver disease, kidney disease, sepsis, and shock. Persisting lactic acidosis may indicate an issue with hepatic metabolism in which the lactate production exceeds the rate the liver can metabolize Persisting high blood lactate concentration is associated with poor prognosis in patient mortality. Measuring blood lactate accurately enables clinicians with an early-prediction prognosis. High levels of local lactate within head, neck and uterine tumor cells may be associated with a greater risk of cancer metastasis. Measuring lactate levels may lead to differentiating between metastatic and benign tumors in those regions. Additionally, early lactate detection may lead to early diagnosis of anastomotic leaks, which are the most feared complication of gastrointestinal surgery which usually involves division of the gastrointestinal tract and removal of a segment of that tract. Early detection and management may minimize morbidity and mortality due to such leaks.

[0006] In clinical practice, blood lactate levels can be measured using central laboratory test equipment, point-of-care blood gas analyzers, spectrophotometric analyzers, and emerging hand-held devices. The conventional analyzers have certain disadvantages, such as the large size, heavy weight, lack of deformability, and complexity of operation. Techniques, such as indirect spectrophotometry, liquid chromatography, and magnetic spectroscopy, require the use of sophisticated equipment and the systems could be expansive, which are unsuitable for wearable, implantable, or disposable applications. With the increasing needs for portable devices, varieties of techniques have been presented, but these techniques often lack sensitivity.

[0007] Thus, redox sensors, including lactate enzyme sensors, having improved sensitivity are desirable. Particularly, in clinical medicine settings, and meeting the increasing demands for improved "point-of-care" portable devices, e.g., demands relating to wearability, implantability, and disposability of the devices, are desirable.

[0008] Use of a metal oxide in a sensor has been disclosed in the prior art. See, for example, U.S. Patent No. 9, 163,313 and U.S. Patent No. 8,552,730, which are both incorporated by reference herein in their entireties. However, these sensors do not comprise a redox catalyst. Enzyme sensors, e.g., for lactate, are also known, but no known enzyme sensors incorporate a redox catalyst, e.g., lactate oxidase, and a metal oxide layer. Exemplary lactate enzyme sensors are disclosed in, for example, U.S. Patent No. 8,999,126, which is incorporated herein by reference in its entirety.

SUMMARY

[0009] Disclosed herein are redox sensors having improved sensitivity due at least in part to the incorporation of a patterned or continuous metal oxide layer therein. The redox sensors disclosed herein may also meet the increasing demands for improved "point-of-care" portable devices. For example, the disclosed sensors may incorporate a flexible substrate, e.g., of paper or Mylar®, and some sensors may be biodegradable.

[0010] In one aspect, a sensor for detecting an analyte is disclosed herein. The sensor may comprise the following: an electrode; a continuous, amorphous, sol-gel metal oxide layer or a patterned metal oxide layer disposed on the electrode; and a redox catalyst in direct contact with an exterior surface of the metal oxide layer. The redox catalyst is a catalyst for oxidation or reduction of the analyte to be detected.

[0011] In some embodiments, the analyte to be detected is selected from the group consisting of lactate, glutamate, dopamine, glucose, cholesterol, creatine, urea, uric acid, pyruvate, alcohol, bilirubin, ascorbate, phosphate, protein, triglyceride, phenylalanine, tyrosine, lipopolysaccharide, hypoxanthine, and combinations thereof. In some embodiments, the analyte is lactate.

[0012] The metal oxide layer may be at least one of IrOx, TiOx, or ZnOx.

[0013] In some embodiments, the electrode may be formed from gold, platinum, copper, aluminum, silver, Sn0 ₂ , ln ₂ 0 ₃ , W0 ₃ , Ti0 ₂ , graphite or glassy carbon. In some embodiments, the electrode is formed on a substrate. [0014] The substrate may be a flexible or rigid substrate. In embodiments where the substrate is a flexible substrate, the substrate may be formed from plastic, metal, paper, or combinations thereof. In embodiments where the substrate is a rigid substrate, the substrate may be formed from glass, metal, plastic, or a combination thereof.

[0015] The redox catalyst may, in some embodiments, be an oxidation catalyst for oxidation of the analyte. In other embodiments, the redox catalyst may be a reduction catalyst for reduction of the analyte. In still other embodiments, the redox catalyst may be an enzyme for oxidation or reduction of the analyte. The enzyme may be any enzyme having as a substrate at least one of the following: lactate, glutamate, dopamine, glucose, cholesterol, creatine, urea, uric acid, pyruvate, alcohols, bilirubin, ascorbate, phosphate, proteins, triglycerides, phenylalanine, tyrosine, and hypoxanthine. For example, the enzyme may be lactate oxidase. In some embodiments, the sensor disclosed herein may be self-referencing, wired, wireless, or biodegradable.

[0016] In another aspect, a sensing system is described herein. The sensing system comprises a sensor as described herein and a receiver or transceiver configured to receive information from the sensor. In some embodiments, the sensing system may comprise one or more amplifiers.

[0017] In another aspect, a method for using a sensor described herein is disclosed. In some embodiments, the method comprises disposing the sensor of claims 1 or 23 in a chemical or biological environment. The chemical environment may be a food product. The biological environment may be a gut or an organ or tissue. The method may further comprise detecting an electrical current from the sensor and correlating the detected electrical current to the presence, absence, or concentration of the analyte in the biological or chemical environment above a minimum detection threshold.

BRIEF DESCRIPTION OF THE DRAWINGS

[0018] Fig 1(a) and 1(b) depict a sensor and methods of making a sensor according to some embodiments herein.

[0019] Fig. 2(a) is a cyclic voltammogram of bare gold electrodes with different sizes. [0020] Fig. 2(b) is cyclic voltammogram of a sensor according to some embodiments disclosed herein.

[0021] Fig. 3(a) is a CV plot of IrO _x film vs. an Ag/AgCL electrode in 1 x PBS at 300 mV/s.

[0022] Fig. 3(b) is a CV plot of an Au film vs. an Ag/AgCL electrode in 1 x PBS at 300 mV/s.

[0023] Fig 4(a) is a CV plot of an Au electrode and IrO _x modified electrode with a size of 100 x 50 microns in PBS at 300 mV/s.

[0024] Fig. 4(b) is a CV plot of the IrO _x electrode in PB S at 300 m V/s.

[0025] Fig. 5 is a time-current plot for the enzyme coated Au electrode in 1 x PBS with the response to lactate.

[0026] Fig. 6(a) is a graph showing the definition for current overshoot (Γ), current fluctuationg (ΔΙ), and transition time (T ₀ ).

[0027] Fig. 6(b) is a graph showing responsive current values at different time points after adding lactate solution.

[0028] Fig. 7(a) is a graph comparing sensitivity for Au electrodes with sensing areas of 1000 x 1000 microns and 100 x 50 microns.

[0029] Fig. 7(b) is a graph comparing sensitivity for Au and IrO _x electrodes with areas of 500 x 500 microns and 1000 x 1000 microns.

[0030] Fig. 8 is a time current plot for WE (with LO _x , i.e., lactate oxidase) and SE (without LO _x ) in 1 x PBS with responses to lactate, glutamate, and dopamine.

[0031] Fig. 9(a) is a time-current plot for WE (with LO _x ) and SE (without LO _x ) in 1 x PBS with response to lactate and glutamate.

[0032] Fig. 9(b) is a time-current plot for WE (with LO _x ) and SE (without LO _x ) in 1 x PBS with response to dopamine only.

[0033] Fig. 10(a) is a time-current plot for the IrO _x modified WE (with LO _x ) and SE (without LO _x ) in PBS with responses to lactate.

[0034] Fig. 10(b) is a time-current plot for the IrO _x modified WE (with LO _x ) and SE (without LO _x ) in PBS with responses to glutamate and dopamine. [0035] Fig. 11(a) is an SEM photo of the Au film before being loaded with enzyme according to some embodiments herein.

[0036] Fig. 11(b) is an SEM showing the structure of the cured lactate protein on the Au surface according to some embodiments herein.

[0037] Fig. 11(c) is an SEM photo of an Au sensor after being used for three weeks according to some embodiments herein.

[0038] Fig. 11(d) is an SEM photo of the bumpy Au surface after being used for one month according to some embodiments herein.

[0039] Fig. 12 (a) is an SEM photo of an IrO _x film before being loaded with enzymes according to some embodiments herein.

[0040] Fig. 12(b) is an SEM photo showing the structure of a cured lactate protein on the IRO _x surface according to some embodiments herein.

[0041] Fig. 12(c) is an SEM of an IrO _x sensor according to some embodiments herein after being used for three weeks.

[0042] Fig. 12(d) is an SEM photo of the IrO _x surface, without enzyme, after being used for a month.

[0043] Fig. 13 is a graph comparing sensitivity between a flat sensor and a bent sensor according to some embodiments herein.

[0044] Fig. 14 depicts a method for making a sensor according to some embodiments herein.

[0045] Fig. 15 depicts a method for making a sensor according to some embodiments herein.

[0046] Fig. 16 depicts a method for making a sensor according to some embodiments herein.

[0047] Fig. 17 depicts a method for making a sensor according to some embodiments herein.

[0048] Fig. 18 depicts a method for making a sensor according to some embodiments herein.

[0049] Fig. 19 depicts a method for making a sensor according to some embodiments herein. [0050] Fig. 20 depicts a sensor according to some embodiments herein.

[0051] Fig. 21 depicts a sensing system according to some embodiments herein.

[0052] Fig. 22 depicts a sensing system according to some embodiments herein.

[0053] Fig. 23 depicts a sensing system according to some embodiments herein.

[0054] Fig. 24 depicts a method for making a microfluidic channel according to some embodiments herein.

[0055] Fig. 25 depicts a microfluidic channel according to some embodiments herein.

[0056] Fig. 26 is a schematic drawing of a potential control circuit according to some embodiments described herein.

[0057] Fig. 27 is a schematic drawing of a trans-impedance and differential amplifier circuit according to some embodiments described herein.

[0058] Fig. 28 is a schematic drawing of a trans-impedance amplifier.

[0059] Fig. 29 is a schematic drawing of a trans-impedance amplifier schematic with guard trace according to some embodiments described herein.

[0060] Fig. 30 is a hardware block diagram of a sensor and sensing system according to some embodiments described herein.

DETAILED DESCRIPTION

[0061] Embodiments described herein can be understood more readily by reference to the following detailed description, examples, and figures (i.e., "FIGs."). Various sensors, devices, systems, and methods for sensing an analyte are described herein, however, such are not limited to the specific embodiments presented in the detailed description, examples, and figures. It should be recognized that these embodiments are merely illustrative of the principles of the present invention. Numerous modifications and adaptations will be readily apparent to those of skill in the art without departing from the disclosed subject matter.

[0062] All ranges disclosed herein are to be understood to encompass any and all subranges subsumed therein. For example, a stated range of "1.0 to 10.0" should be considered to include any and all subranges beginning with a minimum value of 1.0 or more and ending with a maximum value of 10.0 or less, e.g., 1.0 to 5.3, or 4.7 to 10.0, or 3.6 to 7.9. Further, all ranges disclosed herein are also to be considered to include the end points of the range, unless expressly stated otherwise. For example, a range of "between 5 and 10" or "5 to 10" or "5-10" should generally be considered to include the end points 5 and 10.

[0063] Additionally, in any disclosed embodiment, the terms "substantially," "approximately," and "about" may be substituted with "within [a percentage] of what is specified, where the percentage includes 0.1, 1, 5, and 10 percent.

[0064] The terms "a" and "an" are defined as "one or more" unless this disclosure explicitly requires otherwise. The terms "comprise" (and any form of comprise, such as "comprises" and "comprising"), "have" (and any form of have, such as "has" and "having"), "include" (and any form of include, such as "includes" and "including") and "contain" (and any form of contain, such as "contains" and "containing") are open-ended linking verbs. As a result, a composition or other object that "comprises," "has," "includes" or "contains" one or more elements possesses those one or more elements, but is not limited to possessing only those elements. Likewise, a method that "comprises," "has," "includes" or "contains" one or more steps possesses those one or more steps, but is not limited to possessing only those one or more steps.

[0065] It is further understood that the feature or features of one embodiment may generally be applied to other embodiments, even though not specifically described or illustrated in such other embodiments, unless expressly prohibited by this disclosure or the nature of the relevant embodiments. Likewise, devices, systems, and methods described herein can include any combination of features and/or steps described herein not inconsistent with the objectives of the present disclosure. Numerous modifications and/or adaptations of the sensors, devices, systems, and methods described herein will be readily apparent to those skilled in the art without departing from the present subject matter.

[0066] Sensors

[0067] The sensors for detecting an analyte described herein may be redox sensors, biosensors, enzyme sensors, or combinations thereof. The sensors described herein may comprise, consist of, or consist essentially of the following components: (1) an electrode, (2) a continuous, amorphous, sol-gel metal oxide layer or a patterned metal oxide layer disposed on the electrode; and (3) a redox catalyst in direct contact with an exterior surface of the metal oxide layer. The redox catalyst is a catalyst for oxidation or reduction of the analyte being detected by the sensor.

[0068] In some embodiments, the sensor is a self-referencing sensor. A self-referencing sensor, as understood by those skilled in the art, a self-referencing sensor is one in which the analyte does not undergo a reaction at the electrode surface to produce a detectable current, but instead is acted on only by the redox catalyst, e.g., it is oxidized or reduced by the redox catalyst, to generate a species that is detectable at the electrode. In some embodiments herein, an analyte may be oxidized or reduced by the redox catalyst, and in addition, there may be a redox reaction that occurs at the electrode surface and produces a detectable current or other signal.

[0069] In some embodiments, the sensor described herein is a wired or wireless sensor.

[0070] In some embodiments, the sensor described herein is biodegradable. Biodegradable as used herein means that any or all components of the sensor degrade in vivo or in the environment to non-toxic components. In vivo, non-toxic components are those that can be cleared from the body by ordinary biological processes. In some embodiments, a biodegradable sensor completely or substantially completely degrades in the environment or in vivo over the course of about 90 days or less, about 60 days or less, or about 30 days or less, where the extent of degradation is based on percent mass loss of the sensor, and wherein complete degradation corresponds to 100% mass loss.

[0071] The three components of the sensors disclosed herein— the electrode, the metal oxide layer, and the redox catalyst— are discussed in further detail below.

[0072] (1) Electrode

[0073] In some embodiments, the electrode of the sensor may be a working electrode of the sensor. The material of the electrode is not so limited, but is preferably a conductive material such as a metal. Examples of non-limiting materials that may be used to form the electrode include gold, platinum, copper, aluminum, silver, Sn0 ₂ , ln ₂ 0 ₃ , W0 ₃ , Ti0 ₂ , graphite or glassy carbon. In some preferred embodiments, the electrode is formed from gold, platinum, copper, aluminum, or silver.

[0074] The electrode may be free-standing or not-supported by anything, or in some embodiments, the electrode may be formed on or supported by a substrate, including a flexible or a rigid substrate. A rigid substrate may be one that is not easily bent, and in some instances cracks or breaks upon bending. Exemplary rigid substrates may be made of metal, glass, plastic, or some combination thereof. For example, the rigid substrate may be a plastic substrate coated with a metal layer. A flexible substrate, by contrast, is a substrate that is easily bent and does not crack or break upon bending. Exemplary flexible substrate may include paper, e.g., tattoo paper, metal, e.g., a metal foil, a plastic, e.g., Mylar®. Rigid and flexible substrates may be made of the same material, but if this is the case, the rigid substrate will have a greater thickness. In some preferred embodiments, the substrate is flexible, making the sensor more amenable to use in a wearable or portable device. If the substrate is made of a biodegradable polymer or paper, this is also preferable because the sensors will not necessarily have to be removed from the environment where they are used. They may be left in the environment and will biodegrade there.

[0075] The thickness of the electrode is not so limited, but in some embodiments, the thickness may be from 50 nm to 1,000 nm, from 100 to 800 nm, from 200 to 700 nm, from 300 to 500 nm, or from 350 to 400 nm.

[0076] In some embodiments, in addition to the working electrode, the sensor may include one or more additional working, counter, or reference electrodes.

[0077] (2) Metal Oxide Layer

[0078] In some embodiments, the metal oxide layer is a continuous, amorphous, sol-gel metal oxide layer, which is deposited on the electrode.

[0079] A continuous layer is a layer that is unbroken or a layer that does not contain gaps or cracks in the material forming the layer, on any surface of the layer or within the layer, such that the gaps or cracks would prevent the flow of electrons through the layer. More particularly, a continuous layer described herein includes no cracks or gaps that would cause the measured current flowing through the layer, including during operation of a sensor in a manner described herein, to be zero or virtually zero (e.g., 0-0.1 nanoamperes (nA)). For example, in some embodiments, a continuous layer described herein is a layer that does not contain any gaps or cracks larger than about 1 micron (where this size corresponds to the largest size or average size of the gap or crack in a direction parallel to a direction of desired current flow). Preferably, the layer does not contain any gaps or cracks larger than about 0.5 microns, larger than about 0.3 microns, or, most preferably, larger than about 0.1 microns. However, it is to be understood that a "pin hole" or other small gap or crack may be present in a continuous layer described herein, provided the pin hole, gap, or crack does not prevent charge from flowing through the layer in a manner described herein (e.g., through an underlying electrode and into a circuit).

[0080] A continuous layer described herein may be created by any method not inconsistent with the stated goals herein. For example, one method of forming a continuous metal oxide (e.g., IrO _x ) layer without cracks or gaps is disclosed in U.S. Patent No. 9, 163,313, which is incorporated by reference herein in its entirety.

[0081] An amorphous layer, as understood by one of ordinary skill in the art, is noncrystalline.

[0082] The sol-gel metal oxide layer of a sensor described herein can be formed by any sol -gel process not inconsistent with the objectives of the present disclosure. Sol -gel processes are well known and understood by those skilled in the art. Sol-gel methods for forming a metal oxide film comprise, consist of, or consist essentially of, generally, dispersing a metal compound in an appropriate solvent to form a sol, coating the sol on a substrate, and drying the coated sol to remove solvent. An appropriate compound may be a metal alkoxide or a metal halide such as a metal chloride. An appropriate solvent may be water or a polar organic solvent. In some embodiments, dispersing a metal compound in the solvent may comprise dissolving the metal compound in the solvent, but this is not required. The metal compound may be slightly soluble or insoluble in the solvent. The metal compound may be dispersed in the solvent by mechanical means such as mixing. Moreover, coating of the sol may be accomplished by any of the following non-limiting techniques: spin-coating, brush-coating, casting, and dip-coating. Drying may be accomplished, in some embodiments, by exposing the coated sol to temperatures from 20°C to 400°C to remove solvent. It has surprisingly been discovered by the inventors that the use of a sol-gel metal oxide layer provides improved sensor performance, as opposed to the use of a metal oxide layer formed by another method and having a concomitantly differing structure (e.g., a metal oxide layer formed by electroplating or sputtering). These sol-gel metal oxide films can also eliminate gas evolution, pH fluctuation, and metal corrosion side effects, which are common in metal films such as gold.

[0083] Without wishing to be bound by any particular theory, it is believe that this improved sensor performance is due at least in part to the increased surface roughness, of the metal-oxide layers formed by a sol-gel method. Increased roughness means increased surface area and places for the redox catalysts to attach. If the same metal oxide materials are sputtered or electroplated, this same surface structure or microstructure is not observed. The surface is smooth, not rough. Thus, the continuous, amorphous, sol-gel metal oxide layers described herein are not sputtered, electroplated, or otherwise formed by another process other than a sol-gel method because with those processes a rough surface is not obtained.

[0084] The metal oxide of the continuous, amorphous, sol-gel metal oxide layer is not so limited. Any metal oxide capable of forming, using a sol-gel method, a layer that is both continuous and amorphous may be used. Examples of acceptable metal oxides include IrO _x , TiO _x , and ZnOx, with IrO _x being preferred because it has a higher surface charge, which improves the detection limit of the sensor.

[0085] In other embodiments, the metal oxide layer may be a patterned metal oxide layer disposed on the electrode.

[0086] The patterned layer comprises positive space (or areas where the metal oxide material is present) and negative space (or areas where the metal oxide material is not present). The method for forming the patterned metal oxide layer is not so limited. Any method capable of forming a patterned metal oxide layer may be employed. For example, any known photolithographic methods may be used to form the patterned metal oxide layer. For example, a metal oxide layer may be formed, a photoresist mask may be formed on top of the layer, and then the metal oxide layer may be etched and the photoresist mask removed to reveal the patterned layer. Another exemplary photolithographic method is described in Example 2B herein.

[0087] In some embodiments, the patterned metal oxide layer may comprise, consist of, or consist essentially of vertically aligned metal oxide nanotubes. Vertically aligned means that the nanotubes form an angle of between 85 to 95 degrees with the electrode. IrOx, TiO _x , The metal oxide of patterned metal oxide layer is not so limited. Examples of acceptable metal oxides include IrO _x , TiO _x , and ZnO _x .

[0088] The thickness of the continuous, amorphous, sol-gel metal oxide layer or the patterned metal oxide layer is not so limited and may be from 10 to 500 microns, from 20 to 450 microns, from 30 to 400 microns, from 40 to 350 microns, from 50 to 300 microns, from 60 to 250 microns, from 70 to 200 microns, from 80 to 150 microns of from 90 to 100 microns. [0089] (3) Redox Catalyst

[0090] As understood by one of ordinary skill in the art, a redox catalyst is a substance that increases the rate of an oxidation or a reduction reaction without being consumed by the reaction or undergoing any permanent or irreversible chemical change itself. For example, a redox catalyst may increase the rate of an oxidation or reduction reaction by lowering the activation energy of that reaction.

[0091] It is to be understood that the redox catalyst of a sensor described herein is a redox catalyst for one or more specific analytes to be detected by the sensor, as opposed to being a redox catalyst in general, or a redox catalyst for the oxidation or reduction of some other species differing from the analyte of interest. Any redox catalyst not inconsistent with the objectives of the present disclosure may be used. For example, in some embodiments, the redox catalyst is a metal or metal-containing compound or complex. In some such cases, the redox catalyst comprises cerium oxide (e.g., so called "nanoceria"), copper, or a copper N2S2 macrocycle. In other embodiments, the redox catalyst may be a naturally occurring or engineered enzyme. In such cases, the analyte can be a substrate of the enzyme. Moreover, a specific enzyme may have more than one substrate, or chemical reactant, that binds with the enzyme. These substrates may bind individually to the enzyme, or more than one substrate may bind to the enzyme at the same time.

[0092] The redox catalyst is in direct contact with or touching an exterior surface of the metal oxide layer of the sensor. The exterior surface of the metal oxide layer may include a surface of one or more of the pores that may be present on the exterior surface of the metal oxide layer. Further, in some cases, the redox catalyst is attached, bonded, or coupled to the exterior surface of a metal oxide layer described herein.

[0093] In some preferred embodiments, the redox catalyst may be immobilized, physically or chemically, on an exterior surface of the metal oxide layer. Immobilized catalysts are not easily removed from the surface of the metal oxide layer. Methods for physically or chemically immobilizing the redox catalyst include any method known by those skilled in the art. An example of physically immobilizing the redox catalyst on an exterior surface of the metal oxide layer may comprise forming a mixed solution comprising the catalyst and a polymer, and coating this mixed solution on the metal oxide layer. In other embodiments, the catalyst may be physically adsorbed onto an exterior surface of the metal oxide layer. In still other embodiments, the catalyst may be physically trapped or stuck in the pores of the metal oxide layer. Chemical immobilization may involve chemical adsorption of the catalyst onto an exterior surface of the metal oxide layer. In some embodiments, chemical adsorption may involve direct covalent bonding between the catalyst and the metal oxide layer.

[0094] Exemplary methods of binding the redox catalyst are also shown in Fig. 31. These methods include chemically bonding the enzyme to the cross-link material (1 and 2) or adding a physical barrier, e.g., a Nafion semi-membrane, on top of the enzyme (3 and 4).

[0095] Examples of analytes that may be detected by the sensor are not so limited, and may be any analyte with a corresponding redox catalyst capable of oxidizing or reducing the analyte. For example, the analyte may be selected from the group consisting of analyte is selected from the group consisting of lactate, glutamate, dopamine, glucose, cholesterol, creatine, urea, uric acid, pyruvate, alcohols ,e.g., methanol, ethanol, phenols, glycerol, sorbitol, etc., bilirubin, ascorbate, phosphate, protein, triglyceride, phenylalanine, tyrosine, lipopolysaccharide, hypoxanthine, and combinations thereof. In some preferred embodiments, the analyte is lactate.

[0096] In the sensors described herein, the analyte comes into contact with, touches, or in some embodiments binds or selectively binds with the redox catalyst. Subsequently, the analyte is oxidized or reduced by the redox catalyst. Oxidation or reduction of the analyte results in a flow of electrons to or from the analyte. If the analyte is reduced, electrons flow to the analyte. Electrons are a reactant. If the analyte is oxidized, electrons flow from the analyte. Electrons are a product of the oxidation. In some embodiments, electrons flow from the redox catalyst to the electrode, which may act as a transducer. This flow of electrons is translated into an electrical signal by the sensors described herein.

[0097] Systems

[0098] The sensing systems disclosed herein are not so limited. In some embodiments, the sensing systems may comprise, consist of, or consist essentially of at least one sensor as described hereinabove and a receiver, transceiver, or transceiver-receiver configured to receive information from the sensor. In some embodiments, the system may further comprise at least on amplifier. [0099] In some embodiments, different sensors for detecting different analytes may be used in addition to the sensors described herein. In other embodiments, the sensors described herein may themselves have the ability to detect different analytes. For example, the redox catalyst may catalyze the oxidation of more than one different analyte, or the sensors may include a different redox catalyst capable of catalyzing the oxidation or reduction reaction of a different analyte.

[0100] As understood by those skilled in the art, a transceiver is a device comprising both a transmitter and a receiver that are combined and share common circuitry (wires) or a single housing. When no circuitry (no wires) is common between transmit and receive functions, the device is a transmitter-receiver. A receiver is a device that receives a signal and converts the signal into audio or visual form. The signal may be an electrical signal or waves.

[0101] The transceiver, receiver, or transmitter-receiver is configured to receive information from the sensor. This information may be received via wires or wirelessly.

[0102] Methods

[0103] The sensing methods disclosed herein are not so limited and may comprise, consist of, or consist essentially of disposing a sensor as described herein into a chemical or biological environment. Disposing is not so limited. It is only necessary that, upon disposing the sensor into a chemical or biological environment, the analyte, if present in the environment, is able to come in to contact with, touch, bind, or selectively bind with the redox catalyst. In most if not all properly configured sensors, this requirement will be met. Disposing, may in some embodiments involve implanting the sensor into a human or other mammal or animal, dipping the sensor into a solution that may or may not contain the analyte to be sensed, or depositing a solution that may or may not contain the analyte to be sensed onto the electrode comprising the metal oxide layer and the redox catalyst thereon.

[0104] The biological environment is not so limited and may be an in vivo environment. For example, an in vivo environment may be a human gut (or stomach or intestines) or inside any other organ or tissue of a human or mammal. A biological environment may also be a sample of bodily fluid such as blood, urine, saliva, from a human or other mammal. In some embodiments, the sensor may be disposed in sweat. [0105] The chemical environment is also not limited. The chemical environment may be a product, by-product, or intermediate product of an industrial process such as food processing, water treatment, or oil refining. The chemical environment may be a food product such as a dairy product.

[0106] In some embodiments, the method may further comprise, consist of, or consist essentially of the further steps of detecting an electrical current from the sensor and correlating the detected current to a presence, absence, or concentration of the analyte in the chemical environment.

[0107] It is to be understood that "detecting an electrical current," as described herein, can refer to a positive detection event (in which a current is positively detected) or to a negative detection event (in which a current is searched for but not detected over the course of a desired detection window, such a time bin of up to 5 seconds, up to 10 seconds, up to 30 seconds, up to 1 minute, up to 5 minutes, or up to 10 minutes). Moreover, in the case of a positive detection event, detecting the electrical current can indicate the presence of the analyte, and/or can be used to determine an amount or concentration of the analyte (including a non-zero amount or concentration) within the biological or chemical environment. Similarly, in the case of a negative detection event, "detecting" the absence of an electrical current can indicate the absence of the analyte within the biological or chemical environment. Detecting an electrical current may be achieved or carried out in any manner not inconsistent with the objectives of the present disclosure. For example, in some cases, detecting an electrical current is carried out or achieved according to the same principles used in amperometric electrochemical sensors. These devices may include, in some embodiments, a reference electrode in addition to the working electrode. It is to be understood that the working electrode is the electrode including the redox catalyst. Amperometric chemical sensors continuously measure current resulting from a redox reaction. Typically, the current is measured at a constant electrical potential difference or voltage.

[0108] "Correlating," as used herein, does not necessarily refer to mathematical correlation, such as mathematical correlation of variables. Instead, "correlating" refers to using the detected current to identify the presence, absence, or concentration of the analyte. It is generally to be understood that a correlating step can include, without limitation, using a measured amount of current resulting from the redox reaction of the analyte to determine whether an analyte is present or absent in an environment, or to determine the concentration of an analyte in an environment described herein.

[0109] For the correlation step, it is important to understand the relationship between analyte concentration and the number of electrons produced or consumed by the redox reaction at the redox catalyst. For, example it is important to understand whether 2 electrons (or moles of electrons) are produced due to oxidation of one molecule (or mole) of analyte or whether this relationship is 1 : 1, 3 : 1, 4: 1, etc.

[0110] Correlating the detected current to a concentration of the analyte in the biological environment may involve using a calibration curve to match the detected current to a concentration of analyte. A calibration curve may be created by using solutions of known analyte concentration, detecting the electrical current generated, and recording this information in graphical format. A calibrated sensor may be able to make this correlation itself.

[0111] A "minimum detection threshold," for reference purposes herein, is the minimum analyte concentration capable of producing an electrical current that can be detected. This minimum threshold depends on the sensitivity of the sensor. For example, the minimum threshold may be affected by variables such as the number of redox catalyst species present in the sensor or the likelihood that the analyte will interact with the redox catalyst and be oxidized or reduced. When the redox catalyst is an enzyme, the likelihood that the analyte (substrate) will interact with the redox catalyst (enzyme) and be oxidized or reduced may be characterized as enzyme specificity. In some cases, the minimum detection threshold of a sensor, system, or method described herein is above 1 microM, or from 5 to 100 microM, from 10 to 100 microM, from 20 to 100 microM, or from 50 to 100 microM.

EXAMPLES

[0112] Non-limiting Examples are included herein below:

Example 1: Forming a Lactate Sensor

[0113] In one fabrication process, e.g., as shown in Fig. 1(a), the microelectrode was fabricated on a flexible polyimide substrate with a thickness of 125 μπι. Acetone was used to rinse the substrate and the substrate was then heated at 100° C for 15-minutes. E-beam was used for the metal deposition. As an adhesion layer, 50-nm thick chromium layer was deposited before a 200-nm thick gold layer. Photolithography was used to pattern the electrodes different sizes (e.g., 1000 μιη x 1000 μιη ² , 500 μιη x 500 μιη ² , and 100 μιη x 50 μιη ² ). The positive photoresist was spin-coated on the film. Substrate was exposed and developed to form the mask for the electrodes. Gold etchant and chromium etchant were used to etch out the [gold and chromium electrode patterns, respectively.] DI water was used to clean the sample. Acetone was used to clean the remaining photoresist. For testing purposes, silver epoxy connected copper wire to the sensor and then heated at 100°C for 10 minutes to solidify.

[0114] Lactate oxidase (Pediococcus species), L-glutamic salt, and 3-hydroxytyramine hydrochloride were used (Sigma-Aldrich, St. Louis, MO, USA). L-lactic acid (lithium salt) 99% was obtained from Fisher Scientific. 1.3 mg lactate oxidaze ("LOx") was dissolved in 500 μL 1 x phosphate buffer (PBS) solution to form the lactate oxidase enzyme stock solution. Slight agitation was needed to accelerate the dissolution. Then the stock solution was aliquoted to 20 and stored at -20°C. 960 mg L-lactic acid (lithium salt was dissolved in 10 ml 10 x PBS to form 1 M L-lactate stock solution. The stock solution was stored at room temperature. Thirty- eight milligrams of 3-hydroxytyramine hydrochloride was dissolved in 200 mL deionized (DI) water to form the 1 mM dopamine (DA) stock solution. L-glutamic salt (37.428 mg) was dissolved in 200 mL DI water to form the 1-mM glutamate (Glu) stock solution.

[0115] To load the enzymes, the frozen LOx stock solution was kept at room temperature for half an hour until completely thawed. The stock solution was gently agitated by a syringe tip to restore uniformity. The electrodes were cleaned by DI water and dried by air. Ten microliters of stock solution was transferred by a Hamilton syringe and deposited onto the electrode under a stereomicroscope. The same enzyme coating process was repeated four times. The sensors were sealed in a container and kept at the room temperature for two days before tests. During this period of time the protein was cured completely on the metal surface, which prevented the enzyme from dissolving in the solution during further experiments. Hence, the lifetime of the sensors in the treating buffer solution were prolonged.

[0116] The operation of lactate oxidase and the destruction of hydrogen peroxide at the anode of the sensor are based on the equations (1) to (3) below: L-Lartate + C¾ > Pym af & +¾¾ (1)

HA÷ 8(¾ ^* > K ^÷ -i- e- (2)

ΗΟ;· ^* ^ Q _: - * 5Γ - ϊ (3)

[0117] Microfluidic channels may be added to the sensors prior to enzyme injection. This method is illustrated in Fig. 24. A negative photo resist ( PR)(1) is patterned to form a patented mold(2). Then a PDMS layer is formed over the mold (3) and a patented PDMS microfluidic channel is formed (4). Fig. 25 shows a top view (5) and a side view (6) of the PDMS microfluidic channel.

[0118] Results and Discussion

[0119] CV Characterization on Au Electrode

[0120] Cyclic voltammetry was performed on Au electrodes with different sizes. The CV experiments were conducted in the 40 mL PBS, 150 mL KCl solution. Fig. 2(a) shows the current-potential (I-V) curves of Au sensors with the sizes of 1000 x 1000 micrometers and 100 x 500 micrometers. It is obvious that with a larger size the I-V curves were broader, indicating higher current values. Thus, the electrode with the size of 1000 x 1000 micrometers was expected to have a better performance than the smaller size. The CV tests were also performed on the 100 x 500 micrometer Au electrode before and after enzyme coating.

[0121] Fig. 2 shows cyclic voltammograms of (2a) bare gold electrodes with the sizes of 1000 x 1000 micron and 100 x 50 micron; and (2b) the sensor with a size of 100 x 50 micron in PBS with 150 mM KCl before and after enzyme coating. Example 2A: Forming a Lactate Sensor with a Continuous Amorphous Sol-Gel Metal

Oxide Layer

[0122] In Example 2 A, a metal oxide layer was added to the Lactate sensor to improve sensitivity. Specifically, in Example 2A, a lactate sensor with a continuous amorphous sol-gel metal oxide layer is described.

[0123] An exemplary method for forming a lactate sensor with a continuous amorphous sol-gel metal oxide layer is shown in Fig. 1(b).

[0124] An additional metal oxide layer such as an IrO _x thin film deposition was applied. IrO _x has a higher roughness factor compared with a gold seed layer. The expanded surface area was increased and more enzymes were loaded on the IrO _x sensing surface, which increases sensitivity. Confined deposition on patterned gold film formed the IrO _x sensing surface.

[0125] A negative photoresist ( PR) layer SU8-100 (MicroChem) was spin-coated on the gold film as mask. The NPR layer then formed a micro well on the gold surface. During coating, the thicker the NPR layer, the more solution it could capture and stock. However, the increase of thickness can cause the layer become too easy to be peeled off. A sol-gel process was conducted by a process in which 1 g of iridium chloride was dissolved with 42 mL of ethanol and 10 mL of acetic acid to mix the solution. Dip coating was then used to form a 100 micron thin film on the patterned film surface and then it was heated at 75°C for 20 minutes to remove the moisture. The film was then peeled off. A thermal treatment was conducted to the sample with a heating profile starting from 25°C to 325°C for a 3 hour period. The temperature maintained at 325° for 4 hours, then cooled down in a 7-hour period to 25°C. The IrO _x film was then formed. Another layer of NPR (SU8-25) was used afterward as an insulation layer. The copper wire was connected, and the enzyme was loaded on the IrO _x sensing film as described before in Example 1.

[0126] Results and Discussion

[0127] C V Characterization of IrO _x Electrodes

[0128] Increasing the surface roughness to increase the reaction area should lead to an improvement in sensitivity. Cyclic voltammetry was applied to quantitatively analyze the surface roughness of Au and IrOx sensing films. The CV experiments were conducted on the 100 x 50 micron Au and IrOx sensors in the potential window of -0.5 V to +1.0 V in 40 mL 1 x PBS with 150 mM KC1. A scanning rate of 300 mV/s was applied. The roughness factor can be calculated as the ratio of the active reaction area to the geometric area. The equations are shown below: p = A _r /A _g (4)

In (4) and (5) above, p is the roughness factor, A _r is the active reaction area, and A _g is the geometric area. Q _H is the total charge, which can be calculated by taking the integral of the CV curve. Q _H * is the charge density for the single layer molecules of the substrate surfaces. Based on the literature, IrOx has more charge density than Au. In the experiments, Q _H * (the absorption of oxygen) of 386 μθ/αη ² is applied to Au. For IrOx, it is in the range of 500-1900 μθ/αη ² , depending on the film conditions. By taking the integral in the CV plots, the Q _H was obtained. Fig. 3(a) and 3(b) show the integral regions for IrOx and Au (grey area), respectively.

[0129] Figs. 3(a) and 3(b) depicts CV plots of (3a) IrO _x film versus an Ag/AgCl electrode in 1 x PBS at 300 mV/s and (3b) Au film under the same condition. Grey areas were used for calculation of integral regions for Au and IrO _x .

[0130] After calculation, the roughness factor, p, for Au was obtained as 0.512. For IrO _x , Q _H by calculation was 5.69 with Q _H * in the range of 500-1900 μθονη ² , the p, was in the range of 1.2 to 4.5496. Clearly IrO _x had a higher roughness factor and more enzymes could be loaded on the sensing surface. Thus, the IrO _x sensing film increased the sensor sensitivity. Fig. 4(a) shows the CV plots for Au and IrO _x sensors with the same electrode size of 100 x 50 microns in PBS at 300 mV/s. The result showed the conductivity was greatly improved for the IrO _x electrode, with a small reduction peak observed at around 0.19 V.

[0131] CV on Titration Tests

[0132] Titration tests were conducted with the cyclic voltammetry. Fig. 4(b) shows the CV traces for the IrOx sensor before adding lactate (curve #1) and with eight successive accessions of lactate (curves #2-#9). Each time, 80 JL lactate stock solution was applied, which led to concentrations of lactate in the beaker increasing from 2 mM to 16 mM. The oxidation peak current at the bias of approximately -0.05 V increased with respect to each addition of lactate, due to the generation of H ₂ O ₂ in the enzymatic reaction. The reduction peak at approximately 0.15 V also increased, which was caused by the consequential electrocatalytic reduction of H ₂ 0 ₂ . This phenomenon was only observed on the electrode with LOx enzymes.

[0133] Sensitivity Tests

[0134] Chronoamperometry of titration tests was first conducted by the Au sensor with a size of 1000 ^χ 1000 μπι. A constant potential of 0.6 V was applied between the WE and Ag/AgCl reference electrode. The performance of the sensors is shown in Fig. 5. Fig. 5 is a time- current plot for the enzyme coated Au electrode in 1 x PBS with the response to lactate addition. Each addition is 2 mM. The arrows indicate successive addition of lactate solution. Successive additions of 2 mM lactate led to corresponding stepwise increases of electrical currents. After 11 additions, a shot of 320 μΐ _^ lactate was added to the beaker to ensure that the current increases were from the lactate additions. Then another 11 additions of 80 μΐ _^ lactate were added successively. The increased currents produced at the Au anode in the H202 oxidation process were proportional to the lactate concentrations. During lactate additions, the lactate mixture was dripped closely near the gold sensing surface, and then defused into the buffer solution. The lactate concentrations at the certain dipping time points were much higher until they were diluted in the solution. Hence, overshoots of signals were observed when the lactate was first added. To calibrate the sensor and make a consistent discussion on the stability and sensitivity of electrodes based on different sensing films, performance related terms are defined in Fig. 6(a). The red curve in Fig. 6(a) imitated the general current change for one addition of lactate. The units for the graph are relative, which will be defined by the researcher in the experiments. The current overshoot phenomenon and the tendency of current transition were presented. The current overshoot (/') was defined as the difference between the peak current value and 90% of the saturated current value. The overshoots ranged from 0.6-1.5 nA in different additions of lactate. The current fluctuation (ΔΙ) was defined as the current variation range after the sensor reached a stable condition. The current fluctuation may be caused by the system noises such as the electrical noises, electromagnetic interferences, vibration of the testing instrument, and liquid dynamics. The current fluctuation was typically less than 0.15 nA. The transition time (To) was defined as the time period from the beginning of the current overshoot until the current reached 90% of the saturated current value. To investigate the transition time of the sensor after lactate has been added, the current values were measured at different time points after each addition of lactate.

[0135] The results were shown in Fig. 6(b). The value of the x-axis indicates the number of times that lactate was added. The annotation "Xth s" means the time period from the time lactate was added to the buffer solution until the current value was measured. The current values were taken after the lactate was added to the PBS for 10, 20, 50, 100, 130, 150, 160, and 200 s. Based on the results, after 100 s the sensors showed the same current value for different time points, as the current data points overlapped after the 100th s. Hence, it is concluded that the transition time for the Au electrode was 100s. The same experiment was also conducted for the IrO _x electrodes, which showed the same result as the Au electrode. With the measured currents at the 100th s after the overshoots, the titration test showed a sensitivity of 129.6 pA/mM. The titration test was also conducted on the 100 x 50 micrometers Au electrode to investigate the sensitivity for different sensing area sizes. Fig. 7(a) shows the sensitivity comparison between the 100 x 50 micrometers and 1000 x 1000 micrometers Au electrodes. For each sensor size up to 20 electrodes were tested to calculate the average sensitivity. By increasing the sensing area, the average sensitivity increased from 47.5 pA/mM to 129.6 pA/mM. Hence, the sensitivity was improved. However the surface area normalized sensitivity dropped from 950 nA/(cm2 mM) to 13 nA/(cm2 mM). This may be due to the surface tension from the enzyme stock solution on the electrodes. The enzyme mixture was a suspension in which the LOx biomacromolecules were not evenly distributed. The ionic strength of the PBS was interfering with the solubility of the enzyme. After the air-dry process, the proteins most likely located on either the center or the boundary of the solution drop. For the smaller sensing pad, relatively more in terms of percentage of the proteins, were accumulated on the metal. Thus, more current density was produced in a smaller area. This issue may be resolved with robotic suspension to apply the enzyme, which is commonly performed in pharmaceutical practice.

[0136] To increase the surface area normalized sensitivity, the electrode surface was modified with IrOx. IrOx has a higher roughness factor than gold, which makes it possible to accumulate more enzyme proteins. The sensitivities were compared between Au and IrOx modified electrodes with different sizes. For the electrode with the size of 100 x 50 micrometers, the sensitivity increased from 47.5 pA/mM to 462.5 pA/mM, the normalized sensitivity increased from 950 nA/(cm ² mM) to 9250 nA/(cm ² mM). For the electrode with the size of 1000 x 1000 micrometers, the sensitivity increased from 129.6 pA/mM to 1 125 pA/mM , the normalized sensitivity increased from 13 nA/(cm2 mM) to 1 12.5 nA/(cm ² mM). The results of the sensitivity comparison were shown in Fig. 7(b). IrOx increased the surface area normalized sensitivity by 9.17 times for the same sensing size. The surface tension issue of enzyme coating remained as the smaller surface area yields higher normalized sensitivity.

[0137] Selectivity Tests

[0138] Glutamate (Glu) and dopamine (DA) were used as the interference bio-molecules. They were applied to the 1000 x 1000 micrometers Au lactate sensor individually. Fig. 8 shows the sensor current responses to lactate, glutamate, and dopamine. Fifty microliters Glu and 10 \L DA were added in turns after three accessions of 80 \L lactate. The baseline currents were different for the WE (with LOx) and SE (without LOx) because the loaded enzyme changed the impedance of the working electrode (WE). The baseline currents were recalibrated. Then the SE values were subtracted from those of WE to remove the interference effects which were more noticeable for the DA (on the right side of the green dotted curve). The net values (blue dashed curve) showed that the sensor had no responses to Glu and DA.

[0139] The sensor was removed from the beaker and cleaned by 1 x PBS solution. The second and third experiments were conducted separately with respect to Glu and DA. Fig. 9(a) shows the sensor responses to the additions of lactate (L) and glutamate (Glu). The LOx enzyme modified WE had corresponding responses to lactate, while the bare Au SE showed no response. Both WE and SE showed no responses to Glu since it was not an electrode reactive component. The two overshoot signals from the SE were induced by the electron turbulence when the Glu was first added to the beaker. The noises in WE were noticeable compared with those in SE. The reason may due to the interference induced by the chemical reaction conducted on the sensing film. Fig. 9(b) shows the sensor response to dopamine (DA). Both WE and SE showed similar responses caused by the oxidation potential of DA on electrodes. The subtracted values (green dotted curve) showed little response to DA. However, some overshoot and disturbance signals were observed, which happened at the time point when the dopamine was added to the buffer solution. Since the working electrode was covered by the lactate enzyme protein while the self- referencing electrode was directly in contact with the dopamine, there is a response time difference between the two electrodes. Additionally, it was difficult to add the DA solution at the exactly equal distances to the two electrodes. Hence, noise was generated when subtraction for the current responses was done. However after the two electrodes became stabilized, the noise of subtraction reduced. After three additions of dopamine, the current value was still at around 0 nA, same as the initial condition. In the entire time range the subtraction values showed no current increases with respect to the additions of dopamine. Hence, the self-referencing technique can eliminate the interference caused by DA. The selectivity test demonstrated that the sensor probe was responsive only to the additions of lactate.

[0140] The self-referencing technique was also applied to the IrOx -modified sensor. Fig. 10(a) shows the current values conducted with a 1000 x 1000 μπι IrOx sensor. Similar to the Au sensor, the IrOx modified sensor showed corresponding stepwise increased currents with respect to the additions of lactate. The current step each time was approximately 10 times larger than that of the Au electrode with the same size. Same as the Au electrode, the IrOx modified sensor showed no response to interferences, such as glutamate and dopamine, as shown in Fig. 10(b).

[0141] Longevity Tests

[0142] After a few days of use, the Au sensor showed decayed performance. The sensor eventually stopped responding to lactate. This may be due to the loss of weakly-bonded enzymes, the inactivation of the enzyme, or the damage of the sensing film surface. To examine the electrode lifetime in a dry condition, the sensor was first tested in one beaker with four additions of lactate, and then sealed in container for a week at room temperature. The same experiment was conducted and repeated every week. In this case the sensor showed responses to lactate for four continuous weeks. At the fifth week, the sensor started to show a degraded response with less sensitivity. Scanning electron microscopy (SEM) was conducted to check the enzyme quantity and sensing surface quality before and after the sensor was used.

[0143] Fig 11 shows the SEM images of the Au sensing film before and after use. Fig 11(a) shows the condition of Au film before enzyme was loaded. The entire surface was flat and smooth. The small bumps were caused by the dust particle on the polyimide film before metal deposition. Fig. 11(b) shows the structure of the enzyme protein. It clearly shows the protein was evenly distributed on the flat film. Fig. 11(c) shows the Au sensing surface after being used for a month. The amount of enzyme protein decreased compared with Fig. 11(b). This was caused by the dissolution of protein in the PBS solution during the experiment. Fig. 11(d) was taken after the Au electrode stopped working. A bumpy Au surface was observed. It was likely that, after several tests, some of the protein particles were washed away by the buffer solution along with the attached Au film, which left micro-scale pores on the metal layer. The buffer solution leaked through the pores and went under the film to create the bumps. Thus, the sensing pad was damaged.

[0144] Fig. 11(a) SEM photo of the Au film before being loaded with enzymes. Fig. 11(b) SEM photo showing the structure of the cured lactate protein on the Au surface. Fig. 11(c) SEM photo of the Au sensor after being used for three weeks. The amount of the enzyme protein was decreased. Fig. 11(d) SEM photo of the bumpy Au surface after the probe was used for a month.

[0145] Figs. 12(a)-(d) shows the SEM images of the IrOx sensing film before and after use. Fig 12(a) shows the IrOx sensing surface before being loaded with the enzyme. The bumps indicate the cracks of the IrOx crystal which were generated during the heating process. Fig 12(b) and Fig. 12(c) show the structure and distribution of the enzyme protein on the IrOx sensing surface before and after being used for a month. Compared with Fig. 11(c), more of the enzymes were preserved on the IrOx surface after use. Fig. 12(d) was taken after the enzymes were dissolved. The damage of the Au film was not observed for IrOx. Hence, in addition to the increased sensitivity, the rough surface of IrOx allowed better loading of the enzyme and could possibly eliminate gas evolution from the reaction that caused metal corrosion.

[0146] Fig. 12(a) SEM photos of IrOx film before loaded with enzymes. Fig. 12(b) SEM photo showing the structure of the cured lactate protein on the IrOx surface. Fig. 12(c) SEM photo of the IrOx sensor after being used for three weeks. The opening area in the center shows the missing enzyme protein. Fig. 12(d) SEM photo of the IrOx surface after the probe was used for a month.

[0147] Flexibility Test

[0148] The lactate sensors are designed suitable for wearables and implants, owing to the flexibility of the substrate. The sensitivity was tested on a 1000 x 1000 micron sensor in flat and bent conditions. The polyimide substrate supporting the electrodes was bent with a curvature radius of 2 mm. A cotton wire was used to tie the probe shaft to keep the sensor in the bent condition. Fig. 13 shows the comparison between the sensor in the bent and flat conditions. The result demonstrated that the sensitivity was not affected when the sensor was deformed to a curvature radius of 2 mm. A longer response time was observed during experiments. This was likely due to that the sensing electrode facing inwards in the bent probe, hence the applied lactate was not directly touching the electrode. Time for diffusion was needed before the reaction occurred.

[0149] Conclusions

[0150] A lactate-oxidase-based flexible lactate sensor was developed. Two types of biocompatible electrode films, gold and IrOx, as the primary materials for wearable or implantables have been demonstrated. Sensors with different sizes and materials were compared for sensitivity, selectivity, stability, and durability. The rough surface of IrOx provides an improvement in sensitivity. The self-reference technique reduces interference and noise, providing a better selectivity. The simple fabrication method without high thermal budgets provides potentially cost-efficient fabrication of sensors. The flexible polyimide substrates, along with the IrOx electrode being inert, enables the device better biocompatibility for animal and human use. The good performance of the sensing electrodes and the simple fabrication method make an affordable device possible for a variety of practical applications. Disposable devices could be achieved for clinical medicine, food processing, athlete training, and other lactate- detection-related applications.

Example 2B: Forming a Lactate Sensor with a Patterned Metal Oxide Layer

[0151] In Example 2B, a metal oxide layer was added to the lactate sensor to improve sensitivity. Specifically, in Example 2B, a lactate sensor with a patterned metal oxide layer is described.

[0152] IrO _x can be formed in a patterned manner, such as nanotubes, instead of randomly-arranged rough thin film to increase the surface area. An exemplary method is shown in Fig. 19. A thick oxide layer was deposited to form an insulation layer (1). Then, a 2.2 μπι thick negative photoresist ( PR) was spin-coated and patterned to create the bases of microelectrodes and connection lines (2). E-beam evaporation was used to deposit a titanium layer, then a gold layer. The aluminum layer was sputtered with a thickness of at least 900 nm (3). The sacrificial photoresist layer was then removed (4). Another thick layer of PR was patterned as a passivation layer (5). To grow the IrO _x nanotubes, a supportive nanoporous aluminum oxide (AAO) template was needed. An anodization process was conducted, with .23 M oxalic acid solution, to form the AAO. Then, the phosphoric acid 10 wt.% solution was dripped on the AAO template to widen the pore size and remove the aluminum oxide barrier layer at the bottom (5). The IrO _x was then anodoically electrodeposited into the nanoporous structures (6). The sacrificial AAO template was dissolved in potassium hydroxide for 2 hours after (7). Finally, the enzyme stock solution was dropped on the free-standing IrO _x nanotubes and air dried (8).

Example 3: Forming a Lactate Sensor with a Metal a Continuous Amorphous Sol-Gel

Metal Oxide Layer and a Paper Substrate

[0153] In this Example, a sensor was formed by the following steps:

[0154] Step 1 comprises dispensing IrO _x electrode material as described in Example 2 above onto a polymer substrate or a rigid substrate such as silicon or metal. The substrate has electrode shaped molds. See Fig. 14.

[0155] In Step 2, the substrate is placed in an oven or on a hot plate. Temperature increases from room temperature to 325°C or above in 3 hours. Stay at 325°C for 4 hours. Then decrease the temperature back to the room temperature in 7 hours. The IrCl is turned to IrOx.

[0156] In Step 3, silver ink is dispensed on a paper substrate to "print" metal patterns, e.g., to print an electrode pattern and a connection line pattern. See Fig. 15.

[0157] In step 4, the polymer substrate or a rigid substrate with the electrode-shaped IrOx is aligned with the silver electrode pattern so that the two are touching each other. See Fig. 16.

[0158] In step 5, the laminate formed in step 4 is placed on a hot plate at 90°C until the ink is dry. Then the polymer substrate or a rigid substrate is removed leaving the electrode- shaped IrO _x on top of the silver electrode pattern. This step can also be done by photonic curing. See Fig. 16

[0159] Step 6 comprises laminating a thin polymer film or another layer of paper to protect the connection line (Ag). Contact window exposing the IrO _x remain. See Fig. 17. [0160] In step 7, an enzyme (e.g., lactate oxidase) is loaded onto the exposed IrO _x . See Fig. 18

Example 4: Uses of Lactate Sensors In Vivo

[0161] Fig. 20 shows a sensor according to some embodiments herein. The lactate sensor may be made according to the methods described in the Examples above.

[0162] The sensor of Fig. 20 may be used in vivo as shown in Fig. 21, in some sensing systems, the sensor is attached to a catheter and inserted into an abdominal cavity via a port. The sensor is then used to detect lactate in the abdomen of a patient that has undergone gastrointestinal surgery. Lactate detection may lead to early diagnosis of anastomotic leaks, which are the most feared complication of gastrointestinal surgery that usually involves division of the gastrointestinal tract and removal of a segment of that tract.

[0163] Fig. 22 depicts additional configurations of the sensors and sensing systems disclosed herein. These are wired sensors/sy stems, where the sensor is connected to the Rx/Tx box by wires.

[0164] Fig. 23 depicts additional configurations of the sensors and sensing systems disclosed herein. These are wireless sensors/sy stems, where the sensor and Rx/Tx box can communicate wirelessly.

Example 5: Portable Biopotentiostat

[0165] The most important function of a potentiostat is to control the potential at the counter electrode (CE) to allow for current to be measured at the working electrodes (WEI, WE2) due the redox reactions occurring from the applied voltage which is held constant the reference electrode (RE). This is shown by the first op-amp (IC4A) and the functionality was discussed in the principles of a potentiostat section. A second op-amp (IC4B) must serve as a unity gain buffer to limit any current that might otherwise flow through RE. This is necessary because the summing resistors typically have low resistance to minimize thermal noise meaning that significant current would pass through RE if not for this amplifier. These two amplifiers would then need a large open loop gain, small voltage noise density, and small input offset voltages (for potential accuracy). Another factor to consider is that the amplifiers must have a sufficiently high bandwidth (in MHz) to accommodate for the scanning rates of 10, 50, 100, 500, and lOOOmV/s. The LMP7702 (Texas Instruments) op-amp was chosen out of three previously used op-amps for low cost potentiostat design because it has the least noise, and largest open loop gain. The properties of LMP7702 are contained in Table 1 below

[0166] A potential control circuit for a portable bi-potentiostat module is shown in Fig.

26

[0167] Table 1:

[0168] A current measurement circuit or trans-impedance and differential amplifier circuit is shown in Fig. 27. The majority of analog to digital converters (ADC) cannot measure a current directly; the current must first be converted to voltage to obtain a digital signal suitable for recording by a computer. A simple technique that was explored was the current shunt, where the voltage drop is measured across a resistor placed in series between the counter electrode and the control amplifier output. However, this would limit the compliance voltage since the output voltage is not referenced to ground. Therefore, to prevent these limitations trans-impedance amplifiers were employed in series with the working electrode. The inspiration of this symmetric trans-impedance design comes from Nguyen et al. bi-potentiostat design [C. Nguyen et al. "A Wearable System for highly selective L-glutamate neurotransmitter sensing". Topical Conference on Biomedical Wireless Technologies, Networks and Sensing Systems. IEEE Radio and Wireless Week. San Diego, CA. January 2015.]. These must be internally compensated, have small input bias current, and polarization errors. The part chosen was the LMP7721 (Texas Instruments) since it had the lowest input bias current, and greatest gain. The gain is equal to the value of the resistor R _F (100Ω), and C _F was chosen so that the low-pass filter has a cut-off frequency of 34Hz to eliminate the common 60Hz background noises. G= 100V/A

[0169] A basic schematic of a trans-impedance amplifier is shown in Fig. 28. Once the currents are converted to voltages, the differential amplifier is used to remove the common noise from the signal, by subtracting the voltages coming off of both working electrodes. The differential amplifier needs to be able to function at a low voltage supply, have low input voltage noise, and give a sizeable gain. The AD8223 (Analog Devices) was chosen because of its low cost, and single supply voltage range, while still outputting a gain of up to 1000V/V. In order to ensure that only positive values are inputted to the ADC, single supply operation needed to be used but for the INA118 both voltage inputs needed to be 0.98V above ground for linear operation. Specifications for the trans-impedance amplifier LMP7721 are shown in Table 2 below and the specifications for AD8223 are shown in Table 3.

[0170] Table 2- trans-impedance amplifier decision matrix

[0171] Table 3-instrumentation amplifier decision matrix

Input bias current 25nA

[0172] A trans-impedance amplifier schematic with guard trace is shown in Fig. 29. The LMP7721 op-amps are able to pick up a range of up to pico-amps with guard traces, for our design guard traces were not necessary since our range was in μΑβ. The design also filters out noise with the low pass filter and differential amplifier circuit as previously mentioned. However, if space was not limited on the PCB board for future designs, for a smaller current range a guard may be used to surround the input trace and feedback circuitry. It is held at a potential equal to the average input signal potential so that there is no leakage current between them. It also acts as a low-impedance node so that leakage current flows into the guard rather than the input itself, and cancels the effect of added stray and cable capacitance at low frequencies. RG was set to 80.6Ω for a gain of 1000V/V

Example 6: Description of Operation of Sensors

[0173] The electrodes are set-up into a beaker containing the solution of interest, the counter electrode (CE) needs to be fully submerged into the solution, whereas the reference electrode (RE) and the working electrode (WE) need to be halfway submerged. An external Digital to Analog converter is connected to the microcontroller and with an op-amp (OPA277) it produces the cyclic voltammetry in the form of a triangular wave. The negative feedback of the Op-amps (LMP7702) maintain the constant reference potential at the reference electrode (RE), and thus creates a current at the counter electrode (CE). The working electrodes (WEI, WE2) detect the sub μΑ current and pass the signal through trans-impedance amplifiers (LMP7721) which convert the current to voltage. The difference between these two signals is then amplified to eliminate the common noise by passing it through the differential amplifier (AD8223). The signals are then passed to the microcontroller for processing by its internal analog to digital converter, and the data is transferred by USB. A hardware block diagram consistent with the above-description is shown in Fig 30.